Deadtime correction in a nuclear medicine imaging system

ABSTRACT

A method of correcting for deadtime in an attenuation map generated by a nuclear medicine imaging system is provided. The imaging system includes a gamma ray detector that is rotatable about an object to be imaged. A region is defined on the imaging surface of the detector, such as the edge of the field of view of the detector, such that during a transmission scan of the object, the radiation shadow of the object is substantially less likely to fall upon the defined region than upon other regions of the imaging surface. A blank transmission scan is then performed, including recording the radiation intensity level detected in the region as a reference intensity level. A transmission scan of the object is then performed to acquire an attenuation map of the object, including recording a radiation intensity level detected in the first region during the transmission scan. A correction value is then determined as the ratio of the reference intensity level to the intensity level detected in the region during the transmission scan. The attenuation map is then corrected for deadtime based on the correction value.

FIELD OF THE INVENTION

The present invention pertains to the field of medical imaging. Moreparticularly, the present invention relates to deadtime correction in anuclear medicine imaging system.

BACKGROUND OF THE INVENTION

In nuclear medicine, images of internal structures or functions of thebody are acquired by using gamma cameras to detect radiation emitted bya radio-pharmaceutical that has been injected into the patient's body. Acomputer system controls the gamma camera to acquire data and thenprocesses the acquired data to generate the images. Nuclear medicineimaging techniques include single-photon emission computed tomography(SPECT) and positron emission tomography (PET). SPECT imaging is basedon the detection of individual gamma rays emitted from the body, whilePET imaging is based on the detection of gamma ray pairs emitted incoincidence in opposite directions due to electron-positronannihilations. Accordingly, PET imaging is sometimes referred to ascoincidence imaging.

One factor which has an impact on image quality in nuclear medicine isnon-uniform attenuation. Non-uniform attenuation refers to theattenuation of radiation within the body before the radiation can bedetected. Such attenuation tends to degrade image quality. A techniquewhich has been used to correct for non-uniform attenuation istransmission scanning. In transmission scanning, gamma radiation istransmitted through the patient to a scintillation detector and used toform a transmission image, i.e., an attenuation map. The attenuation mapprovides an indication of the amount of attenuation caused by variousstructures of the body and can, therefore, be used to correct forattenuation in the emission images.

Another factor which affects image quality is deadtime loss. Deadtimerelates to the inability of a gamma camera system (the detector, theelectronics, or both) to distinguish between two radiation-inducedscintillation events occurring very close together in time. Deadtimeloss can be defined as the difference between the true countrate (the"singles rate") and the measured countrate due to deadtime. In an idealsystem, in which there is no deadtime loss, the measured countrate wouldequal the true countrate. In contrast, in a real system, which issubject to deadtime loss, the measured countrate is lower than the truecountrate.

Some deadtime correction techniques rely upon approximations of deadtimelosses and are therefore inherently subject to inaccuracies. Inparticular, deadtime losses are dependent, at least in part, upon thesingles rate; as the singles rate increases, deadtime losses tend toincrease. In transmission studies, because the singles rate variesspatially due to many factors, such as the size and shape of thepatient, deadtime losses also tend to be spatially dependent. Correctiontechniques which rely upon approximations typically do not account forsuch variations. Another known technique for correcting for deadtimeloss is to apply a single, global correction factor to the acquireddata, assuming the detector is operating in a limited range of singlesrates. The use of a global correction factor, however, has the samedisadvantages as mentioned above, i.e., it fails to take intoconsideration the spatial dependency of deadtime losses. Other deadtimecorrection techniques rely upon precise calibration of the gamma camerasystem. However, calibration may be time consuming, subject toinaccuracies, and based on incorrect assumptions. Moreover, calibrationalso does not take into account differences in size and shape betweenpatients. Hence, it is desirable to provide a deadtime correctiontechnique which overcomes these and other disadvantages of the priorart.

SUMMARY OF THE INVENTION

The present invention includes a method and apparatus for correcting fordeadtime in an imaging system which uses a radiation detector to acquiredata of an object. In the method, a reference radiation intensity levelrepresenting a blank transmission scan is determined. A transmissionscan of the object is performed to acquire a transmission map of theobject. The transmission scan includes a determination of a radiationintensity level in a defined subset of the imaging surface of thedetector. The transmission map is then corrected for deadtime based on arelationship between the reference intensity level and the intensitylevel detected in the defined subset during the transmission scan.

Another aspect of the present invention is a method and apparatus forgenerating an image in an imaging system such as mentioned above, inwhich a scan of an object is first performed to acquire a data set byacquiring data with the detector at a plurality of angular positionsabout the object. A defined portion of the data set is then used tocorrect the data set. The defined portion of the data set represents aportion of the object about which the detector has acquired data from anumber of pairs of angular positions during the transmission scan, eachof which includes two angular positions approximately 180 degrees apart.

Other features of the present invention will be apparent from theaccompanying drawings and from the detailed description which follows.

BRIEF DESCRIPTION OF THE DRAWINGS

The present invention is illustrated by way of example and notlimitation in the figures of the accompanying drawings, in which likereferences indicate similar elements and in which:

FIG. 1 illustrates a dual-detector gamma camera imaging system.

FIG. 2 illustrates the processing system of the gamma camera system ofFIG. 1, according to one embodiment.

FIG. 3 is a flow diagram illustrating an overall routine for generatingattenuation-corrected coincidence emission images of a patient.

FIG. 4 illustrates a point source transmitting a fan beam to a detectorand a sample area defined on the imaging surface of the detector.

FIGS. 5A, 5B, and 5C illustrate sample areas defined on the imagingsurface of a detector according to three alternative embodiments.

FIG. 6 is a flow diagram illustrating a routine for correcting anattenuation map for deadtime.

FIG. 7 illustrates projection data of an object with a discontinuity inthe data.

FIG. 8 illustrates aspects of a sinogram of transmission data of anobject.

FIG. 9 is a flow diagram illustrating a routine for correctingdiscontinuities in acquired transmission data.

FIG. 10 is a flow diagram illustrating an overall routine for performingrandoms correction.

FIG. 11 is a flow diagram illustrating a routine for acquiring randomsprofiles as a function of object distribution, in accordance with theroutine of FIG. 10.

FIG. 12 is a flow diagram illustrating a routine for acquiring a randomsfraction as a function of both singles and coincidence countrates, inaccordance with the routine of FIG. 10.

FIG. 13 illustrates projection data of an object, including truecoincidences, scatter and random coincidences.

FIG. 14 illustrates three plots of singles rate versus coincidence rate,each corresponding to a different phantom.

FIGS. 15A and 15B are a flow diagram illustrating an alternative routinefor obtaining a randoms fraction.

FIG. 16 illustrates a gamma ray detector fitted with a parallel holecollimator and gamma rays from a Cs-137 point source penetrating thesepta of the collimator.

DETAILED DESCRIPTION

A method and apparatus for performing deadtime correction in a gammacamera imaging system are described. In the following description, forpurposes of explanation, numerous specific details are set forth inorder to provide a thorough understanding of the present invention. Itwill be evident to one skilled in the art, however, that the presentinvention may be practiced without these specific details. In otherinstances, well-known structures and devices are shown in block diagramor other symbolic form in order to facilitate description of the presentinvention.

I. Overview

In certain embodiments, the present invention may be implemented in adual-detector gamma camera system that is capable of both PET and SPECTimaging. Such systems are available from ADAC Laboratories of Milpitas,Calif. FIG. 1 illustrates a block diagram of one such system. Althoughthe illustrated system is a dual-detector system, note that the presentinvention is not limited to implementation in a dual-detector system,nor is the present invention limited to a system that can perform bothPET and SPECT imaging. The gamma camera system of FIG. 1 includes aprocessing system 18 coupled to two scintillation detectors 10. Thedetectors 10 are mounted to a gantry 11, which can rotate the detectors10, either individually or in unison, about an axis of rotation that isperpendicular to the x-y plane (parallel to the z axis). The detectors10 are shown in FIG. 1 configured in a 180 degree orientation relativeto each other about the axis of rotation, such as may be used forcoincidence imaging. A patient 12 rests on a table 14 positioned betweenthe detectors 10. Each of the detectors 10 includes a scintillationcrystal, an array of photomultiplier tubes (PMTs) optically coupled tothe crystal, and appropriate processing circuitry coupled to receive andprocess the outputs of the PMTs and to provide the processed outputs tothe processing system 18.

The gamma camera system also includes two transmission radiation pointsources 16, which are positioned slightly outside the fields of view(FOVs) of the detectors 10. The point sources generally are used fortransmission scanning for purposes of acquiring an attenuation map ofthe patient 12. Each of the point sources 16 is mounted so as to berotatable about the axis of rotation in unison with an opposing detector10. In one embodiment, each of the point sources 16 is a Cesium (Cs-137)662 keV source. Each source 16 is configured to transmit radiation tothe opposing detector in a fan beam profile. The transmission sources 16may be used for transmission imaging for correction of eithercoincidence or SPECT emission data, as will be described further below.

The processing system 18 controls the overall operation of the gammacamera system, including receiving data acquired by the detectors 10,reconstructing images based on the acquired data, controlling the gantry11 to appropriately position the detectors 10, and controlling theimaging mode (PET or SPECT) of the detectors 10 based on a user command.Note that in alternative embodiments, however, some of theabove-mentioned functions or aspects thereof may be implemented withinthe detectors 10, the gantry 11, or in separate modules.

The processing system 18 may be, or may include, a conventional computersystem, such as a personal computer (PC), a server and workstation, asingle-board computer, etc. FIG. 2 illustrates a block diagram ofprocessing system 18 according to one such embodiment. Note, however,that the specific architecture of processing system 18 is not importantfor purposes of practicing the present invention. In the illustratedembodiment, the processing system 18 includes a central processing unit(CPU) 20, random access memory (RAM) 21, read-only memory (ROM) 22, anda mass storage device 23, each coupled to a bus system 28. The bussystem 28 may represent multiple physical buses coupled together byappropriate bridges, controllers, and/or adapters. Also coupled to thebus system 28 are a display device (including controller) 24, which maybe a cathode ray tube (CRT), liquid crystal display (LCD), or the like;a keyboard 25; a pointing device 26, such as a mouse, trackball,touchpad, or the like; a data communication device 27; and a printer 29.Data communication device 27 may be used by processing system 18 tocommunicate with the detectors 10 and/or other computer systems orcomponents and may be, for example, a network adapter, a modem, or anyother suitable data communication device. Display device 24 and printer29 may be used to display and print, respectively, tomographic imagesreconstructed by processing system 18.

It should be noted at this point that some or all aspects of the presentinvention may be embodied in software. That is, the present inventionmay be carried out, at least in part, in a computer system, such asprocessing system 18, in response to its CPU executing sequences ofinstructions contained in memory. The instructions may be executed fromRAM, ROM, a mass storage device, or a combination thereof. In certainembodiments, hardwired circuitry may be used in place of, or incombination with, software instructions to implement the presentinvention. Thus, the present invention is not limited to any specificcombination of hardware circuitry and software, nor to any particularsource of any such software.

It is useful now to briefly consider the overall process associated withimaging a patient. Accordingly, FIG. 3 illustrates an overall routinefor acquiring coincidence emission data and for generating images fromthat data. Note that while the routine of FIG. 3 relates to coincidenceimaging, aspects of the present invention are also applicable to SPECTimaging and associated transmission scans, as will be apparent to thoseskilled in the art. Note also that many variations upon the routine ofFIG. 3 are possible without departing from the scope of the presentinvention.

Referring now to FIG. 3, at 301 a blank (reference) transmission scan isperformed to determine a reference intensity level I₀ for each of theangular positions. A "blank" transmission scan is a transmission scanperformed without a patient or table present and may be performed in apre-clinical test setting. In general, the data from one blank scan canbe used in imaging studies of multiple patients. However, it may bedesirable to perform a new blank transmission scan periodically (e.g.,every few months) to account for drift in system parameters. At 302, ina clinical setting, initial study parameters are set up in processingsystem 18. These parameters may include, for example, the imaging mode(i.e., PET or SPECT, emission or transmission scan, etc.), the initialconfiguration of the detectors, the total number of angular positionsabout the axis of rotation, the total acquisition time at each position,etc. Next, the detectors are configured in the 180 degree orientationrelative to each other about the axis of rotation. Next, at 303,emission and transmission scans of the patient are performed for each ofthe angular positions. Note that the emission and transmission scans maybe performed either simultaneously or sequentially. At 304, processingsystem 18 generates an attenuation map from the transmission data, whichis corrected for deadtime. At 305, the processing system 18 corrects thecoincidence emission data for randoms, and at 306 the processing systemreconstructs the randoms-corrected emission data to generate a set oftomographic images, using the deadtime-corrected attenuation map tocorrect the emission data for non-uniform attenuation.

II. Deadtime Correction

The deadtime correction technique of the present invention requires noknowledge of the singles countrate and requires no a priori knowledgeabout particular detector characteristics. It is well-known that themeasured radiation intensity level I at a depth d within an attenuatingmaterial can be represented as set forth in equation (1):

    I=I.sub.0e.sup.-∫μd                                (1)

where μ is the attenuation coefficient of the attenuating material andI₀ is the measured incident intensity of the radiation upon theattenuating material. From equation (1), we have:

    ∫μd=1n[I.sub.0 /I]                                 (2)

However, equation (2) is not entirely accurate in a system that issubject to deadtime losses. Accordingly, we may rewrite equation (2) fora system which has deadtime losses as equation (3), where D₀ is anappropriate correction factor to correct the blank scan for deadtimelosses, and D is an appropriate correction factor to correct thetransmission scan with the patient present for deadtime losses. ##EQU1##Since I and I₀ are measured values, an attenuation map of the patientcan be corrected for deadtime losses by determining the ratio ##EQU2##

It is useful now to consider the gamma camera system of FIG. 1 ingreater detail. FIG. 4 illustrates one of the detectors 10 and acorresponding one of the point sources 16, with the patient 12positioned between the detector 10 and the point source 16. In oneembodiment, point source 16 is a Cs-137 source configured to transmit662 keV gamma radiation in a fan beam 38, such that the transaxial (x)width of the fan beam 38 at the imaging surface 30 of detector 10essentially matches the transaxial FOV of the detector 10. The width ofthe fan beam 38 is chosen to be relatively narrow in the axial (z)direction. In contrast with certain prior art systems, such as discussedabove, the point source 16 does not move physically across the FOV ofthe corresponding detector 10, but instead remains positioned outsidethe FOV. The fan beam 38 may be scanned axially across the imagingsurface 30 of detector 10 by moving source 16 laterally outside the FOVor by selectively varying collimation (not shown) at source 16.

With the configuration of FIG. 4, a portion 33 of the imaging surface 30will fall within the radiation "shadow" of the patient 12, whereasanother portion 31 of the imaging surface 30 will not fall within theradiation shadow and will therefore receive radiation from source 16 atfull intensity. This is true for each angular position of the detector10. The deadtime correction technique of the present invention is basedupon the assumption that the intensity level I measured in region 31should be identical to the intensity level I₀ measured during thereference scan after both intensity levels have been corrected fordeadtime losses. This assumption, therefore, can be used to determinethe ratio ##EQU3## as will now be described.

In accordance with the present invention, a sample area is defined as aportion 32 of the imaging surface 30 of each detector 10 at the edge ofthe transaxial FOV closest to the source 16. The sample area 32 isselected to be a region of the imaging surface 30 upon which theradiation shadow of the patient is least likely to fall during animaging session, or is substantially less likely to fall than otherregions of the imaging surface 30. Accordingly, let thedeadtime-corrected transmission intensity at a given angle measured insample area 32 during the blank scan be represented as I₀,SA •D₀,SA.Similarly, let the deadtime-corrected transmission intensity at thatangle measured in sample area 32 when the patient is present berepresented as I_(SA) •D_(SA). Thus, the above assumption can berepresented by equation (4).

    I.sub.SA •D.sub.SA =I.sub.0,SA •D.sub.0,SA     (4)

From equation (4 ), we have: ##EQU4##

We can use equation (5) to substitute ##EQU5## for the quantity ##EQU6##in equation (3). Therefore, the present invention provides that theratio of the intensity level I_(SA) in the sample area 32 during apatient scan to the intensity level I₀,SA measured in the sample area 32during the reference scan is used as a deadtime correction factor forcorrecting the transmission map of the patient. This ratio may bedetermined for one angular position and then applied uniformly to theentire transmission map. Alternatively, separate ratios can bedetermined for each angular position and applied individually to thetransmission map. The latter approach is believed to be more accurate.

Note that the above-described technique generally is based on theassumption that deadtime losses are uniform across the imaging surfaceof the detector. Because that assumption may not hold, however, thepresent invention also includes a technique to correct for localvariations in deadtime losses across the imaging surface of a detector,as will be described below.

Refer now to FIGS. 5A through 5C, which illustrate a detector 10 viewedperpendicular to its imaging surface. As shown, the sample area 32 canbe defined along the edge of the transaxial FOV 40 of the detector 10.In the embodiment of FIG. 5A, the sample area 32 is defined along theentire axial field of view of detector 10. As noted, the sample area 32is defined such that the radiation shadow 42 of the patient will notcover (or will rarely cover) the sample area 32. In certain instances,such when scanning an extremely large patient, the radiation shadow may,in fact, fall upon the sample area 32 for certain angular positions.Accordingly, an alternative approach, illustrated in FIG. 5B, is todefine two (or more) separate sample areas 32a and 32b adjacent to eachother axially at the edge of the transaxial FOV of the detector 10.Thus, even if the patient's radiation shadow falls upon one of thesample areas for a given angular position, it is likely that the shadowwill not also fall upon the other sample area at the same time, becausea patient's body contour tends to be irregular. Accordingly, for eachangle, the greater of the intensities measured in each of sample areas32a and 32b is taken as the intensity level ISA for purposes ofcomputing the deadtime correction factor. Numerous other approaches todefining one or more sample areas are possible within the scope of thepresent invention. FIG. 5C, for example, shows yet another embodiment,in which two sample areas 32c and 32d are defined to overlap.

FIG. 6 illustrates a routine for correcting an attenuation map of thepatient according to the above-described technique. More specifically,the routine of FIG. 6 illustrates in greater detail part 304 of theroutine of FIG. 3. Note that the following routine is performedindependently for each detector. Initially, at 601 the average ratio##EQU7## inside the sample area is determined for a given angle θ. Asingle computation of this ratio may be used for all angles, as notedabove. However, the routine of FIG. 6 makes a separate computation foreach angle. Next, at 602 the appropriate attenuation coefficient ∫μd isdetermined according to equations (3) and (5) for each (x,y) position.If there are additional angles to be considered at 603, then the nextangle is selected at 604, and the routine repeats from 601; otherwise,the routine ends. Thus, an attenuation map generated as described aboveis appropriately corrected for deadtime.

As noted above, the foregoing technique assumes that the deadtime isuniform across the imaging surface of the detector. That assumption,however, may not be correct. If the detector deadtime varies along theimaging surface of the detector, then the sampling at the outer edge ofthe transaxial FOV of a detector may not represent the deadtime on theother side of the transaxial FOV, which samples the center of theobject. Small data errors in projection data representing the center ofthe object can cause substantial artifacts in the reconstructed image.

Refer again to FIG. 4, which shows the axis of rotation 36 of thedetectors 10. The axis of rotation 36 is included within a center region34. Center region 34 is defined by the inner edge of the fan beam 38(i.e., the edge toward the right side the FIG. 4 for the illustratedorientation) for all angular positions in aggregate. Note that thecenter region 34 will be sampled by the inner edge of the transaxial FOVof the detector 10. (In certain systems, the center region 34 may becovered by supplementing rotation of the detectors 10 with appropriatemotion of the table 14 to extend the transaxial FOV, if necessary.)However, because the illustrated system is a dual-detector system inwhich the detectors 10 are rotated about the patient, the center region34 can be covered twice by each detector after 360 degrees of rotationof the detector.

Because the center region 34 can be covered twice by each detector, thedata acquired the first time it is covered by a detector should be thesame as the data acquired the second time it is covered by the detector,if there is no variation in deadtime across the imaging surface of thedetector. Referring to FIG. 7, which shows a sample projection, anydiscrepancy between the two samplings of the center region 34 may appearas a discontinuity 50 at the center 51 of the projection. Again, such adiscontinuity may cause substantial artifacts in the reconstructedimage. In accordance with the present invention, such discontinuitiesare corrected generally by averaging values in sinogram space whichrepresent a discontinuity and then recomputing those values based on theaverage. Thus, the discontinuity is smoothed by such a technique,eliminating or reducing the severity of artifacts in the reconstructedimage. This procedure is performed for all projections and for alltransaxial slices. Consequently, this procedure corrects for certainlocal variations in deadtime along the imaging surface of a detector.

FIG. 8 illustrates the format of a sinogram 56 of an object. In the dualrotating detector system described above, each detector acquires dataalong each of a number of diagonal lines 58 during a transmission scan.Each diagonal line 58 on the left side of sinogram 56 represents dataacquired by a detector while the detector is at a particular angularposition, while the corresponding diagonal line 58 on the right side ofsinogram 56 represents data acquired by the same detector whenpositioned at the 180 degrees opposite angular position. Centerline 60of sinogram 56 corresponds to the axis of rotation 36 of the detectors10 (FIG. 4). Similarly, the center strip 63 of the sinogram 56corresponds to center region 34. The diagonal lines 58 of data extendbeyond centerline 60, because the inner edge of the transmission fanbeam 38 extends beyond the axis of rotation 36.

In accordance with one embodiment of the present invention, thediscontinuity correction technique described above is applied to allprojection data. FIG. 9 illustrates a routine embodying this technique.At 901, for a given projection θ, the processing system 18 computes anaverage of the data values within the center strip 63 of the sinogram 56which result from detector rotations defined as left side angles (i.e.,the left side of the sinogram). This average is designated as the leftaverage AVG_(L). Also at 901, the processing system 18 computes anaverage of data values within the center strip 63 of the sinogram 56which result from detector rotations defined as right side angles. Thisaverage is designated as the right average AVG_(R). Note that if onetransaxial slice does not provide enough center strip data to bestatistically significant, multiple transaxial slices can be averagedtogether for purposes of computing these averages.

Since corresponding left and right data values represent the same objectdensity, such data values should be the same. In this context,"corresponding" left and right data values are horizontally-aligned datavalues on each side of the centerline 60 and equidistant from thecenterline 60. For example, point 65 in FIG. 9 represents a left datavalue, while point 66 represents the corresponding right data value. Dueto local deadtime variations, however, corresponding left and right datavalues may not be the same. The result may be a glitch in theprojection, as shown in FIG. 7. Consequently, two correction factors areapplied in accordance with the present invention, to correctly align theprojection.

Specifically, after computing the averages AVG_(L) and AVG_(R), at 902 aleft correction factor F_(L) and a right correction factor F_(R) arecomputed based on the averages. The correction factors F_(L) and F_(R)are computed so as to move toward the average all corresponding left andright data values, respectively, for the given projection θ. Forexample, the left correction factor F_(L) may be computed as F_(L)=(AVG_(R) +AVG_(L))/(2•AVG_(L)), while the right correction factor F_(R)may be computed as F_(R) =(AVG_(R) +AVG_(L))/(2•AVG_(R)).

Next, at 903 a left data value LEFT in the given projection θ ismultiplied by the left correction factor F_(L), and a right data valueRIGHT in the projection θ is multiplied by the right correction factorF_(R). The computation of 903 is then repeated for all data values inthe given projection θ, per 904 and 907 (i.e., for all data values alonga given horizontal line in the sinogram), using the same left and rightcorrection factors for all data values in that projection θ. When allpairs of data values in the projection θ have been adjusted, 901 through904 are repeated for all transaxial slices, per 905 and 908, includingcomputing new left and right averages AVG_(L) and AVG_(R) and newcorrection factors F_(L) and F_(R) for each slice. The foregoing stepsare then further repeated for all other projections θ represented in thesinogram (906 and 909). Thus, projection data are effectively scaledbased on the proposition that two data values representing the samepoint should be the same when covered from either of two angles 180degrees apart. This technique, therefore, corrects for local variationsin deadtime along the imaging surface of a detector.

III. Randoms Correction

The present invention also includes a technique for correctingcoincidence emission data for randoms. One problem with randomscorrection techniques of the prior art is that they fail to take intoconsideration variations in the size and shape of patients. Accordingly,the technique of the present invention uses the actual patientdistribution to perform randoms correction. Further, a basic premise ofthe present invention is that the randoms fraction R_(f) may vary eventhough the singles rates remains constant, due to the distributedradiation source in the object, radiation sources from outside the FOV,the energy window used, and the camera geometry. Hence, randomscorrection is also performed as a function of both the singles rate andthe coincidence rate. More specifically, both the coincidence rate andthe singles rate are used in determining the randoms fraction R_(f).

As will be described further below, patient projection data P(x, y, θ)is corrected based on a computed randoms distribution R(x, y, θ). Therandoms distribution R(x,y,θ) is computed as a function of a measuredrandoms profile R₀ (x, y, θ) and a randoms fraction R_(f) (S,C). Therandoms fraction R_(f) (S,C) is computed as a function of both thesingles rate S and the coincidence rate C for each angle θ. The randomsprofile R₀ (x, y, θ) is computed as the convolution of the patientprofile P(x, y, θ) with a wide (high σ) Gaussian distribution. Thisapproach therefore differs from randoms corrections techniques of theprior art, which do not take into account the coincidence rate or theobject distribution.

Refer now to FIG. 10, which illustrates an overall routine forperforming randoms correction according to the present invention. At1001, the gamma camera system obtains raw coincidence (emission)projection data P(x, y, θ) for all angles θ. Optionally, at 1002, thegamma camera system computes axial averages P(x, y, θ) of multipletransaxial slices. For example, each axial average P(x, y, θ) may be anaverage of three (or any other appropriate number of) axially adjacenttransaxial slices. Such averaging may be omitted, if desired. At 1003,processing system 18 computes the randoms profiles R₀ (x, y, θ). At1004, processing system 18 computes the randoms fraction R_(f) (S, C),where S represents the singles rate for a given angle θ, and Crepresents the coincidence rate for that angle θ. At 1005, processingsystem 18 computes a randoms distribution R(x, y, θ) according toequation (6),

    R(x, y, θ)=R.sub.0 (x, y, θ)•R.sub.f (S, C) (6)

At 1006, the processing system 18 corrects the averaged patientprojection data P(x, y, θ) for randoms by subtracting the randomsdistribution data R(x, y, θ) from the averaged projection data P(x, y,θ) to produce corrected patient projection data P_(c) (x,y,θ), as setforth in equation (7).

    P.sub.c (x, y, θ)=P(x, y, θ)-R(x, y, θ)  (7)

FIG. 11 illustrates a routine for computing the randoms profiles R₀ (x,y, θ). Note that most of the routine of FIG. 11 is intended to beperformed in a pre-clinical test setting, rather than in a clinicalsetting, as will become clear from the description which follows. Theexception is that the final computation of the randoms profile R₀ (x, y,θ) is performed in a clinical setting, i.e., during an imaging session.Thus, at 1101, the gamma camera system is operated to acquire projectiondata for coincidence activity emitted by a simple phantom, for multipleangles θ, at both high countrate and low countrate. A simpleright-cylindrical phantom is believed to be suitable. In this context,"low" countrate is a countrate at which randoms are negligible (e.g.,below one percent of the total coincidence events). Next, at 1102 theacquired high and low countrate datasets are each corrected for systemdeadtime using any suitable deadtime correction technique, such as thetechnique described above. Next, at 1103 randoms R(x,y) are computed asthe difference between the high countrate projection data and the lowcountrate projection data for a given angle θ. At 1104 the σ value ofthe above-mentioned Gaussian function is computed. The Gaussian functionis defined according to equation (8) and is a well-known mathematicalfunction. ##EQU8##

The σ value is computed by minimizing the difference between the randomsdistribution R(x,y) and the convolution of the projection data from thehigh countrate phantom and the Gaussian G(x,y), i.e., by minimizing((P(x,y)×G(x,y))-R(x,y)). A simple curve fit procedure can be employedusing, for example a chi-square or a least squares approach. For animaging system such as described above, a σ value of 200/√2 is currentlybelieved to be appropriate, according to one embodiment.

The final aspect of determining the randoms profile R₀ (x, y, θ) isperformed in the clinical setting, based on the patient projection data.More specifically, at 1105 the randoms profile R₀ (x, y, θ) is computedas the convolution of the axially averaged projection data P(x, y, θ)and the Gaussian distribution G (x,y), using the σ value computed asdescribed above, as set forth in equation (9).

    R.sub.0 (x,y, θ)=P(x,y,θ)×G(x,y)         (9)

FIG. 12 illustrates a routine for computing the randoms fraction R_(f)as a function of both singles rate S and coincidence rate C, accordingto one embodiment. Note that the routine of FIG. 12 is intended to beperformed in a pre-clinical setting. At 1201, the gamma camera system isoperated to obtain coincidence projection emission data for each ofmultiple phantoms at various different countrates. Correspondingcoincidence rates C and singles rates S are recorded during the dataacquisition process. The phantoms are selected to fairly represent awide range of patient body sizes and shapes. In one embodiment, threeright-cylindrical phantoms of varying sizes are used. At 1202, thesingles rates S and corresponding coincidence rates C are plotted foreach phantom. FIG. 14 shows an example of three C vs. S plots 81, 82 and83, corresponding to three different phantoms. At 1203, a randomsdistribution R is determined from the data for each of multiple (C,S)pairs by using a linear extrapolation, as will now be described withreference to FIG. 13.

FIG. 13 illustrates a plot 70 of projection data taken from acylindrical phantom from a given projection angle θ. The overallintensity I (i.e., total number of counts) detected can be representedas the sum of all detected true coincidences T, all detected scattercoincidence Sc, and all detected randoms R. Thus, portion 73 of plot 70represents the contribution of randoms R, portion 72 represents thecontribution of scatter Sc, and portion 71 represents the contributionof true coincidences T. At 1203, above, portion 73 representing thecontribution from randoms is approximated using a linear extrapolation.More specifically, coincidence events are measured at the perimeter ofthe FOV, represented by shaded regions 74 under curve 70. The height ofregions 74 at their innermost boundaries is linearly extrapolated acrossthe face of the detector, as represented by line 76 in FIG. 13. Thus,the area of curve 70 which falls below line 76 approximates the totalnumber of counts due to randoms. Note that this approximation does notaccount for a small fraction of the contribution of randoms, which isrepresented by the shaded area above line 76 under curve 70. However, inanother embodiment, to be discussed below, the randoms fraction R_(f)may be computed in a manner which avoids this type of approximationerror. Further, even if the randoms fraction R_(f) is underestimated,the shape of the randoms profile computed from equation (9) is notaffected by such error.

Thus, at 1203 the randoms R are determined for each of multiple (C, S)pairs using this linear extrapolation. The randoms fraction R_(f) isthen computed by dividing the randoms R by the total number of countsfor each of the (C, S) pairs. The randoms fraction R_(f) is then plottedagainst singles rate S at 1205, and at 1206, the randoms fraction R_(f)is plotted against coincidence rate C for various constant singles ratesS. At 1207, the R_(f) versus C curves are fitted to equation (10) todetermine the coefficients A and B.

    R.sub.f =A•S.sup.2 /C+B                              (10)

In one embodiment of the present invention, the values A=1.09×10⁻⁵1/kcps and B=0.19 are believed to be appropriate for a system such asdescribed above. Note, however, that these coefficients may be differentfor other embodiments due to many factors, such as the specific designof the imaging system to be used.

Finally, at 1208 equation (10) with the computed coefficients A and B isstored in any suitable form within the gamma camera imaging system, foruse during an imaging session. Equation (10) is then applied during theimaging of a patient to compute R_(f) for various different angles θ,based on the measured singles and coincidence rates. More specifically,computed R_(f) values are used to determine the overall randomsdistribution R(x, y, θ) according to equation (6), which is used tocorrect the patient projection data P_(c) (x, y, θ) according toequation (7).

It is believed that, at low countrates, the formula of equation (10) mayovercorrect under count-poor conditions. Accordingly, an alternativeformula for R_(f) may be used when the singles countrate falls below athreshold countrate. The formula of equation (11), for example, may besuitable for countrates below 450 kcps for a gamma camera system such asdescribed above. ##EQU9##

As noted above, the foregoing technique for computing the randomsfraction R_(f) may underestimate the total randoms R as a result of thelinear extrapolation. Therefore, FIGS. 15A and 15B collectivelyillustrate an alternative routine for computing the randoms fractionR_(f), which avoids this approximation and which also may provide a moreaccurate σ value for the Gaussian function G(x,y). At 1501, the gammacamera system is operated to obtain coincidence (emission) projectiondata of a phantom at both high countrate and at low countrate. The datais then plotted. As noted above, low countrate in this context refers toa countrate at which the randoms are negligible, such as 100 kcps orbelow. Next, at 1502 the high and low countrate data are corrected forsystem deadtime. At 1503, a randoms curve is computed as the differencebetween the high countrate projection data and the low countrateprojection data. At 1504, the randoms fraction R_(f) is computed bydividing the total number of counts R under the randoms curve by thetotal number of counts under the high countrate curve Total_(H). At1505, the computed randoms curve is fitted to equation (12) in order todetermine the value of σ. A chi-square or least squares curve fit may beused. Note that the R_(f) values are summed before performing the curvefit.

    R=R.sub.f (X)e.sup.-x.spsp.2/.sup.2σ.spsp.2          (12)

The preceding steps are then repeated for other countrates to obtainadditional σ values (1506). The foregoing steps are then furtherrepeated for other phantoms of different sizes to obtain additional σvalues (1507). Next, at 1508, the average σ value σ is computed for allthe previously determined σ values, and at 1509 the average σ is takento be the final a value of the Gaussian function G(x,y). At 1510, athree-dimensional plot of C vs. S vs. R_(f) is computed from theacquired data, and at 1511 the plot is stored in any suitable form inthe imaging system, for later use during an imaging system (i.e., forcomputation of R_(f)). For example, the plot may be stored in the formof a look-up table.

IV. Dual-Use Transmission Source

As noted above, the transmission point sources 16 of the above-describedimaging system may be Cs-137 sources, which transmit radiation with aphotopeak at 662 keV. A Cs-137 point source has been found to bedesirable for performing transmission scans for correction of PET datain such a system. Because the 662 keV photopeak of Cs-137 is close tothe 511 keV photopeak of flourodeoxyglucose (FDG), a common PETradio-pharmaceutical, but not so close as to cause excessive energylevel overlap, minimal energy level scaling of the transmission map isrequired. Further, because the 662 keV photopeak of Cs-137 is higherthan the FDG photopeak, contamination due to scatter into thetransmission image is reduced. In addition, the 662 keV photopeak ofCs-137 is sufficiently high as to enable the majority of the transmittedgamma rays to completely penetrate most general-purpose nuclear medicinecollimators. This characteristic also makes the use of Cs-137advantageous for performing transmission scans to correct SPECT data, aswill now be discussed.

Gamma ray detectors are typically collimated during SPECT imaging, andparallel-hole collimators are commonly used. FIG. 16 illustrates aparallel-hole collimator 82 coupled to the imaging surface 30 of one ofthe detectors 10. The collimator 82 includes a number of lead septa 81,which define the parallel holes. FIG. 16 further illustrates gamma rays83 from the Cs-137 transmission source 16 penetrating the septa 81 ofcollimator 82 to reach the imaging surface 30 of the detector 10. Asnoted above, certain prior art systems use line sources, which arescanned directly across the FOV of the corresponding detectors. Such atechnique allows a substantial number of the transmitted photons to passthrough the holes of the collimator to the detector, since the paths ofmany of the transmitted photons are parallel to the holes of thecollimator. Therefore, there is no need to remove the collimator for thetransmission scan in such a system.

For various reasons, however, it may desirable to use a transmissionsource configuration in which the sources remain outside the FOV duringthe transmission scan, as in the above-described system (see FIG. 4). Asresult, the transmitted gamma rays will impinge upon the collimator atan acute angle, as shown in FIG. 16. In such a system, if a standardsingles transmission source (i.e., a source with a photopeaksubstantially below 662 keV) were used with a standard SPECT collimator,few (if any) of the transmitted photons would pass through thecollimator to the detector, due to the incident angle of the photonsrelative to the holes of the collimator. Rather, most of the photonswould be absorbed by the septa of the collimator. Removal of thecollimator would therefore be required prior to performing atransmission scan, which can be time-consuming and difficult.

However, by using a 662 keV Cs-137 transmission source in accordancewith the present invention, a high percentage (generally, at least 50%)of the transmitted photons impinging upon the collimator will completelypenetrate the septa to reach the imaging surface when using ageneral-purpose nuclear medicine collimator (except for a 511 keVcollimator). Collimators with energy ratings below 300 keV are believedto be preferable, however, to increase photon penetration. Penetrationof ⁵⁰ % or more is sufficient to generate a high-quality transmissionmap of the patient for correcting SPECT data using current dataprocessing techniques. Thus, by using a Cs-137 transmission source,removal of the collimator is not necessary prior to performing atransmission scan for correcting SPECT data. Note that the reconstructedtransmission map must be scaled appropriately to match the energy levelof the single-photon emission source. Techniques for performing suchscaling are well known to those skilled in the art.

Thus, a method and apparatus for performing deadtime correction in agamma camera imaging system have been described. Although the presentinvention has been described with reference to specific exemplaryembodiments, it will be evident that various modifications and changesmay be made to these embodiments without departing from the broaderspirit and scope of the invention as set forth in the claims.Accordingly, the specification and drawings are to be regarded in anillustrative sense rather than a restrictive sense.

What is claimed is:
 1. A method of correcting for deadtime losses in animaging system having a radiation detector, the radiation detectorhaving an imaging surface, the method comprising:determining a referenceradiation intensity level representing a blank transmission scan;performing a transmission scan of an object to acquire a transmissionmap of the object, including determining a radiation intensity level ina defined subset of the imaging surface of the detector; and correctingthe transmission map for deadtime losses based on a relationship betweenthe reference intensity level and the intensity level detected in thedefined subset during the transmission scan.
 2. A method according toclaim 1, wherein the defined subset is defined such that a radiationshadow of the object is substantially less likely to intersect thedefined subset than other subsets of the imaging surface during thetransmission scan of the object.
 3. A method according to claim 1,wherein the defined subset comprises an edge of a field of view of thedetector.
 4. A method according to claim 1, wherein correcting thetransmission map comprises correcting the transmission map based on theratio of the reference intensity level and the intensity level detectedin the defined subset during the transmission scan of the object.
 5. Amethod according to claim 1, wherein performing the transmission scan ofthe object comprises acquiring transmission data with the detector ateach of a plurality of angular positions about the object anddetermining a radiation intensity level in the defined subset at each ofthe angular positions; and wherein the method comprises:determining areference intensity level for each of the angular positions; and foreach of the angular positions, correcting the transmission map based ona relationship between the reference intensity level for said angularposition and the intensity level detected in the defined subset at saidangular position during the transmission scan of the object.
 6. A methodaccording to claim 1, wherein performing the transmission scan of theobject comprises acquiring transmission data with the detector at aplurality of angular positions about the object, the method furthercomprising:storing data acquired by the detector during the transmissionscan as a sinogram; and correcting for variations in deadtime across theimaging surface of the detector by using a defined portion of thesinogram to correct the data, said portion of the sinogram representinga portion of the object about which the detector has acquired data froma plurality of pairs of angular positions during the transmission scan,each pair including two angular positions approximately 180 degreesapart.
 7. A method according to claim 6, wherein the portion of thesinogram comprises data representing a center of the object, and whereinsaid correcting for variations in deadtime comprises, for each of aplurality angular positions represented in the sinogram, correctingasymmetry in data values representing symmetrical locations about thecenter of the object.
 8. A method according to claim 1, furthercomprising:recording a radiation intensity level in a second definedsubset of the imaging surface of the detector during the blank scan as asecond reference intensity level; and recording a radiation intensitylevel detected in the second defined subset during the transmissionscan; wherein said correcting the transmission map comprises correctingthe transmission map for deadtime losses further based on a relationshipbetween the second reference intensity level and the radiation intensitylevel detected in the second defined subset during the transmissionscan.
 9. A method of correcting for deadtime in a medical imaging systemfor acquiring images of an object, the imaging system having a firstradiation detector, the first radiation detector having an imagingsurface, the method comprising:performing a blank transmission scanusing the detector, including recording a radiation intensity leveldetected in a defined region of the imaging surface of the detector as areference intensity level, the defined region characterized in that,during a transmission scan of the object, a radiation shadow of theobject is substantially less likely to fall upon the defined region thanupon other regions of the imaging surface; performing a transmissionscan of the object using the detector to acquire a transmission map ofthe object, the transmission scan including recording a radiationintensity level detected in the defined region; determining a firstcorrection value representing a relationship between the referenceintensity level and the intensity level detected in the defined regionduring the transmission scan of the object; and correcting thetransmission map for deadtime based on the first correction value.
 10. Amethod according to claim 9, wherein the defined region comprises anedge of the detector.
 11. A method according to claim 9, wherein theblank transmission scan and the transmission scan of the object eachcomprise acquisition of transmission data with the first detector at aplurality of angular positions about the object, and wherein the methodcomprises:determining the first correction value for each of theplurality of angular positions to generate a plurality of firstcorrection values; and for each of the plurality of angular positions,correcting the transmission map for deadtime based on the correspondingfirst correction value.
 12. A method according to claim 9, furthercomprising:defining a second region on the imaging surface of the firstdetector, the second region characterized in that, during a transmissionscan of the object, a radiation shadow of the object is substantiallyless likely to fall upon the second region than other regions of theimaging surface of the first detector; recording a radiation intensitylevel in the second region during the blank scan as a second referenceintensity level; recording a radiation intensity level in the secondregion during the transmission scan; determining a second correctionvalue, the second correction value representing a relationship betweenthe second reference intensity level and the intensity level detected inthe second region during the transmission scan and; selecting either thefirst correction value or the second correction value based on the firstand second correction values; wherein correcting the transmission mapfor deadtime based on the first correction value comprises correctingthe transmission map for deadtime using the selected one of the firstcorrection value and the second correction.
 13. A method according toclaim 9, wherein the transmission scan of the object comprisesacquisition of transmission data with the first detector at a pluralityof angular positions about the object, the method furthercomprising:storing data acquired by the first detector during thetransmission scan of the object in a sinogram; and correcting forvariations in deadtime across the imaging surface of the first detectorby using a portion of the sinogram to reduce discontinuities in thesinogram, the portion representing a portion of the object from whichdata is acquired by the first detector from at least one pair of angularpositions 180 degrees apart during the transmission scan of the object.14. A method according to claim 13, the portion of the sinogramincluding data representing a center of the object.
 15. A method ofcorrecting for deadtime in a gamma camera system for acquiring images ofan object, the gamma camera system having a first gamma ray detector,the first gamma ray detector having an imaging surface, the methodcomprising:a) identifying a first region on the imaging surface of thefirst detector, the first region characterized in that, during atransmission scan of the object, a radiation shadow of the object issubstantially less likely to fall upon the first region than otherregions of the imaging surface of the first detector; b) performing ablank transmission scan using the first detector, including recording aradiation intensity level detected in the first region as a firstreference intensity level; c) performing a transmission scan of theobject using the first detector to acquire a transmission map of theobject, including recording a radiation intensity level detected in thefirst region during the transmission scan; d) determining a firstcorrection value, the first correction value representing a relationshipbetween the first reference intensity level and the intensity leveldetected in the first region during the transmission scan; and e)correcting the transmission map for deadtime based on the firstcorrection value.
 16. A method according to claim 15, wherein the firstregion comprises an edge of the detector.
 17. A method according toclaim 15, wherein the blank transmission scan and the transmission scanof the object each comprise acquisition of transmission data with thefirst detector at a plurality of angular positions about the object, andwherein the method comprises:determining a first correction valueaccording to said step d) for each of the plurality of angularpositions; and correcting the transmission map for deadtime based on thefirst correction value for each of the plurality of angular positions.18. A method according to claim 15, defining a second region on theimaging surface of the first detector, the second region characterizedin that, during a transmission scan of the object, a radiation shadow ofthe object is substantially less likely to fall upon the second regionthan other regions of the imaging surface of the firstdetector;recording a radiation intensity level in the second regionduring the blank scan as a second reference intensity level; recording aradiation intensity level in the second region during the transmissionscan; determining a second correction value, the second correction valuerepresenting a relationship between the second reference intensity leveland the intensity level detected in the second region during thetransmission scan; comparing the first correction value to the secondcorrection value and; selecting either the first correction value or thesecond correction value based on said comparing; wherein correcting thetransmission map for deadtime based on the first correction valuecomprises correcting the transmission map for deadtime using either thefirst correction value or the second correction value, based on saidselection.
 19. A method according to claim 15, wherein the transmissionscan of the object comprises acquisition of transmission data with thefirst detector at a plurality of angular positions about the object, themethod further comprising:binning data acquired by the first detectorduring the transmission scan of the object into a sinogram; andcorrecting for variations in deadtime across the imaging surface of thefirst detector by using a portion of the sinogram to reducediscontinuities in the sinogram, the portion representing a portion ofthe object from which data is acquired by the first detector from atleast one pair of angular positions 180 degrees apart during thetransmission scan of the object.
 20. A method according to claim 19, theportion of the sinogram including data representing a center of theobject.
 21. A method of generating an image of an object using animaging system having a radiation detector, the methodcomprising:performing a scan of the object to acquire a data set byacquiring data with the detector at a plurality of angular positionsabout the object; and using a defined portion of the data set to correctthe data set, said defined portion of the data set representing aportion of the object, such that the detector has acquired the datacorresponding to the portion of the object at a plurality of pairs ofangular positions about the object, such that each of the pairs ofangular positions includes two angular positions of the detector thatare approximately 180 degrees apart.
 22. A method according to claim 21,wherein the data set comprises a sinogram, and wherein the portion ofthe sinogram comprises data representing a center of the object, andwherein said using a defined portion of the data set comprises, for eachof a plurality angular positions represented in the sinogram, correctingasymmetry in data values in the sinogram that represent symmetricallocations about the center of the object.
 23. A method according toclaim 21, wherein the scan is a transmission scan of the object, suchthat the data set is a transmission data set.
 24. An apparatus forcorrecting for deadtime in an imaging system having a radiationdetector, the radiation detector having an imaging surface, theapparatus comprising:means for determining a reference radiationintensity level representing a blank transmission scan; means forperforming a transmission scan of an object to acquire a transmissionmap of the object, including means for determining a radiation intensitylevel in a defined region of the imaging surface of the detector; andmeans for correcting the transmission map for deadtime based on arelationship between the reference intensity level and the intensitylevel detected in the defined region during the transmission scan of theobject.
 25. An apparatus according to claim 24, wherein the definedregion comprises an edge of a field of view of the detector.
 26. Anapparatus according to claim 24, wherein the means for correcting thetransmission map comprises means for correcting the transmission mapbased on the ratio of the reference intensity level and the intensitylevel detected in the defined region during the transmission scan of theobject.
 27. An apparatus according to claim 24, wherein the means forperforming the transmission scan of the object comprises means foracquiring transmission data with the detector at each of a plurality ofangular positions about the object and means for determining a radiationintensity level in the defined region at each of the angular positions;and wherein the apparatus comprises:means for determining a referenceintensity level for each of the angular positions; and means forcorrecting the transmission map for each of the angular positions, basedon a relationship between the reference intensity level for the angularposition and the intensity level detected in the defined region at theangular position during the transmission scan of the object.
 28. Anapparatus according to claim 24, wherein the means for performing thetransmission scan of the object comprises means for acquiringtransmission data with the detector at a plurality of angular positionsabout the object, the apparatus further comprising:means for storingdata acquired by the detector during the transmission scan as asinogram; and means for correcting for variations in deadtime across theimaging surface of the detector by using a defined portion of thesinogram to correct the data, said portion of the sinogram representinga portion of the object about which the detector has acquired data froma plurality of pairs of angular positions during the transmission scan,each pair including two angular positions approximately 180 degreesapart.
 29. A gamma camera system comprising:a gantry; a gamma raydetector mounted to the gantry so as to be movable about an object to beimaged; a radiation transmission source; and a processing system coupledto the gantry, the detector, and the transmission source, the processingsystem configured to control the gantry, the detector, and thetransmission source to:perform a blank transmission scan using thetransmission source and the detector to measure a reference radiationintensity level; define a region of the detector in which radiationreceived from the transmission source during a transmission scan of theobject is not likely to have passed through the object; perform thetransmission scan of the object using the transmission source and thedetector to acquire a transmission map of the object, includingilluminating the object with radiation from the source, and measuring asecond radiation intensity level in the defined region of the detector;and correct data within the transmission map for deadtime based on theratio of the reference intensity level to the second intensity level.30. An apparatus according to claim 29, wherein the defined regioncorresponds to an edge of a field of view of the detector.